Third-generation CT scanners typically include an x-ray source and an array of x-ray detectors secured respectively on diametrically opposite sides of an annular disk, the latter being rotatably mounted within a gantry support. During a scan of a patient located within the opening of the disk, the disk rotates about a rotation axis while x-rays pass from the focal spot of the X-ray source through the patient to the detector system.
The x-ray source and detector array are positioned so that the x-ray paths between the focal spot and each detector all lie in the same plane (the so-called "slice plane", "rotation plane" or "scanning plane") which is normal to the rotation axis of the disk. Because the ray paths originate from substantially a point source and extend at different angles to the detectors, the ray paths resemble a fan, and thus the term "fan beam" is used to describe all of the ray paths at any one instant of time. The radiation that is detected by a single detector at a measuring interval of time during a scan is considered a "ray". The ray is partially attenuated by the mass of the patient in its path and each detector generates a single intensity measurement as a function of the attenuation, and thus of the density of the portion of the patient in the path of the ray from the focal spot to that detector. These x-ray intensity measurements, or projections, are typically performed during prescribed measurement intervals at each of a plurality of angular disk positions.
Ideally, all of the radiation within the fan beam should be of uniform intensity during the scan, and all of the detectors should have a uniform input-output response (or transfer function), i.e., all of the detectors ideally should provide the same output signal for a given input signal level of X-radiation. In addition, ideally there should be no variation in the stability of the response of the detector system to the radiation, i.e., the signals produced from the detector system should not drift between successive or periodic scans.
Various types of detectors have been developed, including gas and solid state types. A typical solid state detector includes a scintillating crystal which converts high energy x-radiation photons into low energy visible light photons, and a photodiode which converts the low energy visible light photons into extremely low-amplitude electrical currents (i.e., on the order of picoamperes to nanoamperes). The extremely low-amplitude current output of each detector represents the x-ray flux incident on the detector. The outputs of the detector array are transmitted via an array of conductors to a data acquisition system (DAS) for signal processing.
Because resolution of the resulting image is a function of the size of the detectors, a CT scanner system typically includes hundreds of detectors which are extremely closely spaced within the fan beam arc. For reducing the costs of such detector arrays, preassembled solid state detector modules, each comprising several solid state detectors, have been described in U.S. Pat. No. 5,487,098 issued to John Dobbs and David Banks, and assigned to the present assignee (hereinafter, "the '098 Patent"), and incorporated by reference into this application. For example, one third-generation CT scanner system manufactured by the present assignee includes 384 detectors provided by 24 modules of 16 detectors each and closely spaced within an arc which subtends not more than 48 degrees. The width of a single detector is thus on the order of a millimeter.
Each detector module is typically enclosed on all sides in an electrically conductive, optically opaque shield which is substantially transparent to x-rays. The shield around the module is typically formed from a thin reflective foil. In addition, each individual detector (hereinafter, "detector" or "detector crystal") within a module is surrounded by a white, highly reflective material which is on the order of at least 0.2 millimeter thick and which permits passage of x-rays, yet prevents excessive light leakage between the crystals. X-rays are thus able to impinge on the detector crystal from any angle; however, visible light generated in the crystal in response those x-rays is unable to pass, or reflect, from one detector to another because of the optically opaque shield around the detector, thus substantially reducing or eliminating cross-talk between adjacent detector channels.
Because of the coating thickness between adjacent detector crystals and the mechanical tolerances associated with the manufacture of the individual crystals, it is impractical to position the crystals sufficiently close to one another to achieve a truly continuous detecting region. Even the most closely spaced detector crystals in an array are separated by a distance of at least about 0.2 mm to accommodate the multiple layers of reflective material between them. This small spacing, or gap, between individual detector crystals is manifested as a region of diminished signal information in each projection of the patient. The existence of such regions diminishes the effectiveness with which the detector crystals can uniformly detect the radiation across the fan beam, and thus influences the accuracy and the resolution of the reconstructed image.
This condition is aggravated by the necessity to collimate the radiation prior to its impingement on the detector crystals. Because dense matter tends to scatter x-rays, it is necessary to preclude, to the greatest extent possible, the impingement upon the detector crystals of any radiation that does not traverse a straight path from the focal spot of the x-ray source to each detector crystal. To preclude the impingement of such stray or scattered radiation on the detector crystals, an array of elongated, thin collimation, or "anti-scatter", plates is positioned between the x-ray source and the detector crystals. The anti-scatter plates are opaque to X-rays and are aligned relative to the detector crystals so as to collimate, and thus permit passage of, substantially only those rays traversing a straight line from the source to a detector crystal. The anti-scatter plates are generally placed so that they are aligned with radial lines extending from the focal spot and corresponding gaps between adjacent detector crystals of the array so that they block any rays that impinge on the detector crystal at an angle which varies, for example, by no more than about three degrees from a normally incident ray along the respective ray path. The alignment of each anti-scatter plate with a corresponding gap between detectors insures that each anti-scatter plate shadows the least sensitive part of adjacent detectors, and thus a maximum amount of radiation is detected by the corresponding detector crystals. An advantageous anti-scatter plate alignment system is also described in the '098 patent.
In existing CT scanner systems, it is assumed that at least some spacing between each detector crystal in an array is necessary to accommodate the anti-scatter plates and the shadows they cast on the detector crystals. While each anti-scatter plate is extremely thin, e.g., on the order of 0.1 mm thick, it still occasionally blocks x-rays or creates a "shadow" on one or both of the corresponding adjacent detector crystals. A detector crystal which is "shadowed" by an anti-scatter plate generates a signal which is correspondingly reduced in strength relative to a signal from an unshadowed detector crystal. However, detector system response is affected much more by signal instability than by reduced signal strength. When the disk carrying the detector crystals and anti-scatter plates rotates during a scan, the anti-scatter plates can move over time as a result of thermal effects or of the rotation of the disk. This movement over time of the anti-scatter plates relative to the detector crystals produces variations in the extent of shadowing by an anti-scatter plate on a corresponding detector crystal. The result of this variable shadowing is that the output signal of a shadowed detector may be modulated by the moving anti-scatter plate. Even though the motion of the anti-scatter plates is slight in terms of spatial movement, because the output of each detector crystal is calibrated to an extremely precise degree (0.03%) and is used to make measurements at this level of precision, the amount of fluctuation in the current amplitude can be significant. Thus, prior to the present invention, it was believed that each anti-scatter plate should be as thin as possible and aligned as closely as possible with the corresponding gap between two adjacent detector crystals (the least sensitive part of each detector crystal) so as to minimize the casting of shadows on the detecting regions of the adjacent detector crystals, and thus minimize the modulation of the output signal from the detectors.
However, the optimum system design requires tradeoffs in the design of the detector assembly. On the one hand, it is desirable to make the gap between detector crystals as small as possible in order to more closely simulate a continuous detecting region. On the other hand, it is desirable to make the gap between adjacent detector crystals large enough to accommodate shadows from the anti-scatter plates and thus minimize the effects of the shadows cast by the plates on the detecting regions. This design tradeoff makes it highly impractical to locate the anti-scatter plates so that their shadows fall entirely within the gaps between the detector crystals and not on the detecting regions of the detector crystals. The sensitivity of a detector crystal to radiation is generally at a constant maximum value over the central detecting region of the detector crystal and falls off sharply near the edges of the crystal. Shadowing of the edge, and therefore a less sensitive region, of a detector crystal, i.e., near the gap between adjacent detector crystals, will still cause significant changes in signal intensity for relatively small movements of the anti-scatter plate, thus producing undesirable signal modulation and instability.
As a result of these conditions, extremely painstaking and accurate placement of the detector modules and the anti-scatter plates relative to one another has been required. To minimize shadowing, maximize the amount of direct radiation that is detected, minimize the detection of stray radiation, and prevent signal instability, virtually all relative movement between the plates and the detector crystals must be minimized or prevented. As the resolution of the detector crystals becomes greater, the crystals and the gaps between them become smaller, and the ability to keep the anti-scatter plate shadows entirely within the gap regions becomes increasingly impractical to achieve and maintain. Further, because the detecting region is not continuous but is instead an array of discrete detecting regions of substantially constant maximum sensitivity separated by regions of decreased sensitivity at the edges of a detector crystal, a small but potentially significant portion of the total radiation passing through the patient may not be detected. The resulting signals from the detectors are likely to be relatively unstable, and the images reconstructed from those signals may be distorted and obscured by artifacts introduced thereby.
It would therefore be an advantage in the art of CT scanner systems to overcome the disadvantages of the present systems.